Multi-energy x-ray source

ABSTRACT

An apparatus for use in a radiation procedure includes a radiation filter having a first portion and a second portion, the first and the second portions forming a layer for filtering radiation impinging thereon, wherein the first portion is made from a first material having a first x-ray filtering characteristic, and the second portion is made from a second material having a second x-ray filtering characteristic. An apparatus for use in a radiation procedure includes a first target material, a second target material, and an accelerator for accelerating particles towards the first target material and the second target material to generate x-rays at a first energy level and a second energy level, respectively.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates generally to systems and methods for imageacquisition and, more specifically, to systems and methods forangiogenesis imaging using computed tomography.

2. Background of the Invention

Each year, many women are diagnosed with breast cancer, and the numberof deaths associated with breast cancer has reached 40,000 per year. Thedeath toll cannot be substantially reduced because current screening anddiagnostic techniques do not detect all cancers at an early enough stageto effectuate a treatment cure. As such, early diagnosis and treatmentof breast cancer is highly desirable.

Tumors larger than a few millimeters in cross section generally requirea blood supply in order to obtain nutrients and oxygen for growth.Vessels that grow around the tumors proliferate in a disorganized mannerand they may leak and pool blood around the tumors.Iodine-contrast-enhanced mammography has been used to detect tumors inwomen's breasts. In such procedure, contrast agent is introduced into apatient's vessel, and mammograms of the patient's breast are obtainedbefore and after the contrast injection. By digitally subtracting thepre-injection image from post-injection images that are obtained over atime period of 1 to 7 minutes, a composite image can be obtained thatshows tumor blood supply and “pooling” in the vicinity of the tumor.However, the contrast resolution of mammography is usually limited, andmalignancies in tissues that are below 5 millimeters in cross-sectionaldimension may not be detectable.

In order to create the composite image, the post-injection imagesobtained over the prescribed time period need to have a same imageregistration with the pre-injection image such that a pixel in thepost-injection images can be processed with a corresponding pixel in thepre-injection image. Sometimes, a portion of the patient being imagedmay move, e.g., translates 1 to 400 microns and/or rotates 1 to 2degrees, when between images are taken. In such cases, a physician wouldneed to realign the two images. If the pre-injection image and thepost-injection images cannot be realigned or registered, the imagecollection procedure will need to be repeated. Usually, in a mammographyprocedure, a patient's breast is compressed by a set of paddles toreduce movements of the breast while pre-injection and post-injectionimages are obtained. However, patients usually experience discomfortfrom compression of the breasts. The discomfort can become so severethat the imaging procedure is terminated. In addition, the use ofpaddles to compress breast tissue restrict blood flow into the breast,thereby limiting an amount of contrast agent that can be delivered intothe breast. This in turn, interferes with the kinetics of the contrastagent, and makes angiogenesis imaging difficult.

For the foregoing, improved systems and methods for imaging angiogenesisare desirable.

SUMMARY OF THE INVENTION

In accordance with an embodiment of the invention, a method ofgenerating images of a portion of a body includes introducing a contrastagent into the body, generating a first set of image data usingradiation at a first energy level after the contrast agent is introducedinto the body, generating a second set of image data using radiation ata second energy level after the contrast agent is introduced into thebody, and creating a composite image using the first and the second setsof image data. In one embodiment, the first energy level is below ak-edge of the contrast agent, and the second energy level is above ak-edge of the contrast agent. In one embodiment, the composite image iscreated by subtracting a logarithmic transform of the first set of imagedata from a logarithmic transform of the second set of image data. Suchtechnique results in tissue features being removed from the image whilefeatures attributable to the introduced contrast agent are retained inthe image.

In accordance with another embodiment of the invention, a system forgenerating CT image data using radiation at a plurality of levels isprovided. The system includes a gantry, a x-ray source assembly securedto the gantry, and a detector assembly. By means of non-limitingexamples, the x-ray source assembly can include a plurality of voltagesupplies, a plurality of target materials (anodes), and/or a pluralityof filters, thereby allowing the x-ray source assembly to deliverradiation having different characteristics. The detector assembly caninclude a single imager for generating image data in response toradiation at different energy levels. Alternatively, the detectorassembly can include a plurality of imagers, each of which configured togenerate image data in response to radiation at a prescribed range ofenergy levels. By means of non-limiting examples, the imager can includea scintillating material that converts x-ray into light, or aphotoconductor layer that produces electron-hole-pairs in response tox-ray radiation.

Other aspects and features of the invention will be evident from readingthe following detailed description of the preferred embodiments, whichare intended to illustrate, not limit, the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The drawings illustrate the design and utility of preferred embodimentsof the present invention, in which similar elements are referred to bycommon reference numerals. In order to better appreciate how advantagesand objects of the present invention are obtained, a more particulardescription of the present invention briefly described above will berendered by reference to specific embodiments thereof, which areillustrated in the accompanying drawings. Understanding that thesedrawings depict only typical embodiments of the invention and are nottherefore to be considered limiting of its scope, the invention will bedescribed and explained with additional specificity and detail throughthe use of the accompanying drawings in which:

FIG. 1 illustrates a computed tomography system in which embodiments ofthe present invention may be implemented;

FIGS. 1A and 1B illustrate variations of the computed tomography systemof FIG. 1;

FIG. 2 is a flow chart illustrating a method for imaging angiogenesis ina tissue;

FIG. 3A is a graph showing a low energy spectrum curve, a high energyspectrum curve, and a k-absorption edge of a contrast agent;

FIG. 3B is a graph showing how concentration of contrast agent variesover time for normal and abnormal tissue;

FIGS. 4A-4F illustrate x-ray source assemblies in accordance withdifferent embodiments of the invention;

FIG. 5 illustrates a detector assembly in accordance with an embodimentof the invention;

FIG. 6 illustrates a variation of the detector assembly of FIG. 5,showing the detector assembly having a photoconductor;

FIG. 7 illustrates an imager in accordance with another embodiment ofthe invention;

FIG. 8 illustrates a detector assembly having a plurality of imagers inaccordance with another embodiment of the invention; and

FIG. 9 is a diagram of a computer hardware system with which embodimentsof the present invention can be implemented.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Various embodiments of the present invention are described hereinafterwith reference to the figures. It should be noted that the figures arenot drawn to scale and elements of similar structures or functions arerepresented by like reference numerals throughout the figures. It shouldalso be noted that the figures are only intended to facilitate thedescription of specific embodiments of the invention. They are notintended as an exhaustive description of the invention or as alimitation on the scope of the invention. In addition, an aspect or afeature described in conjunction with a particular embodiment of thepresent invention is not necessarily limited to that embodiment and canbe practiced in any other embodiments of the present invention.

Referring now to the drawings, in which similar or corresponding partsare identified with the same reference numeral, FIG. 1 illustrates acomputed tomography (CT) image acquisition system 10 in accordance withsome embodiments of the present invention. The system 10 includes agantry 12, and a panel 14 for supporting a patient 16. The gantry 12includes a x-ray source assembly 20 that projects a beam of x-rays, suchas a fan beam or a cone beam, towards a detector assembly 24 on anopposite side of the gantry 12 while a portion of the patient 16 ispositioned between the x-ray source assembly 20 and the detectorassembly 24. In the illustrated embodiment, the x-ray source assembly 20is configured to deliver radiation at a plurality of energy levels, andthe detector assembly 24 is configured to generate image data inresponse to radiation at different energy levels. The x-ray sourceassembly 20 may include a collimator 21 for adjusting a shape of thex-ray beam. In some embodiments, the collimator 21 includes one or morefilters (not shown) for creating radiation with certain prescribedcharacteristics. The detector assembly 24 has a plurality of sensorelements configured for sensing a x-ray that passes through the patient16. Each sensor element generates an electrical signal representative ofan intensity of the x-ray beam as it passes through the patient 16.

In the illustrated embodiment, the panel 14 has an opening 60 throughwhich a breast 18 of the patient 16 can be placed when the patient 16 islying on the panel 14 facing downward. Such arrangement allows thebreast 18 to be placed between the x-ray source assembly 20 and thedetector assembly 24. In an alternative embodiment, the panel 14 can beused to support the patient 16 facing upward (FIG. 1A). In such case,the gantry 12 is placed above the panel 14, and is configured to rotateabout the breast 18. In another embodiment, the gantry 12 is configuredto rotate about the breast 18 while the patient 16 is standing (orsitting) in an upright position (FIG. 1B). It should be noted that thepositioning of the gantry 12 should not be limited to the examplesillustrated previously, and that the gantry 12 can have otherconfigurations (e.g., positions or orientations of an axis of rotation),depending on a position of the panel 14, and a position and orientationof an object for which imaging is desired.

In the illustrated embodiment, the CT image acquisition system 10 alsoincludes a processor 54, a monitor 56 for displaying data, and an inputdevice 58, such as a keyboard or a mouse, for inputting data. Theprocessor 54 is coupled to a control 40. The rotation of the gantry 12and the operation of the x-ray source assembly 20 are controlled by thecontrol 40, which provides power and timing signals to the x-ray sourceassembly 20 and controls a rotational speed and position of the gantry12 based on signals received from the processor 54. The control 40 alsocontrols an operation of the detector assembly 24. For example, thecontrol 40 can control a timing of when image signal/data are read outfrom the detector assembly 24, and/or a manner (e.g., by rows orcolumns) in which image signal/data are read out from the detectorassembly 24. Although the control 40 is shown as a separate componentfrom the gantry 12 and the processor 54, in alternative embodiments, thecontrol 40 can be a part of the gantry 12 or the processor 54.

During a scan to acquire x-ray projection data (i.e., CT image data),the x-ray source assembly 20 projects a beam of x-rays towards thedetector assembly 24 on an opposite side of the gantry 12, while thegantry 12 rotates about the breast 18. In one embodiment, the gantry 12makes a 360° rotation around the breast 18 during image dataacquisition. Alternatively, if a full cone detector is used, the system10 may acquire data while the gantry 12 rotates 180° plus the angle ofthe beam pattern. Other angles of rotation may also be used, dependingon the particular system being employed. In one embodiment, the detectorassembly 24 is configured to generate at least 900 frames of images inless than 1 second. In such case, the gantry 12 only needs to rotatearound the breast 18 once in order to collect sufficient amount of imagedata for reconstruction of computed tomography images. In otherembodiments, the detector 24 may be configured to generate frames atother speeds.

FIG. 2 is a flowchart 200 illustrating a method for imaging angiogenesisin the breast 18 using the system 10. First, a contrast agent isintroduced inside the patient's body, and more specifically, into avascular system of the patient 16 (Step 202). The contrast agent can beadministered with a mechanical power injector via an intravenouscatheter that is placed in antecubital or forearm vein, at a ratebetween 2 to 6 milliliter (mL) per second. In the illustratedembodiment, about 50 to 70 mL of contrast agent is administered to thepatient 16. However, other amounts of contrast agent can be introducedinside the patient's body, depending on a patient's size, and/or arequirement of a particular procedure.

A variety of contrast agent can be administered to the patient 16. Inthe illustrated embodiment, the contrast agent includes iodine, whichhas a k-absorption edge (K-edge) of 33 keV. Alternatively, gadolinium(Gd) (having a k-edge of 50.2 keV) chelated withdiethylenetriaminepentaacetic acid (DTPA) can be used. Gd-DTPA is welltolerated by humans, and no serious side effects have been reported. Thecontrast agent can also include materials, such as holmium (having ak-edge of 56 keV), erbium (having a k-edge of 58 key), lanthanum,cerium, praseodymium, neodymium, samarium, europium, terbium,dysprosium, thulium, ytterbium, lutetium (having a k-edge of 63.3 keV),and other rare earth elements. Compounds, DTPA complexes,ethylenediamine tetraacetic acid (EDTA) complexes, nitrilotriacetic acid(NTA) complexes, and other chelate compounds, formed from any of theabove mentioned elements can also be used. Elements with atomic numbershigher than that of gadolinium is particularly suitable for the contrastagent because x-ray absorption of body tissue would be lower at higherx-ray photon energies. However, elements with atomic numbers lower thanthat of gadolinium can also be used, depending on a particular k-edgerequirement of an application. K-edge energies for various materials areknown. Other soluble, non-toxic, chelate compounds can also be used. Inaddition, noble gases such as Xenon, and agents composed of stableisotopes of radio nuclides such as Ti, Yb, Cs, Xe, I, In, and Tc.

Next, the patient 16 is positioned such that the breast 18 of thepatient 16 is positioned between the x-ray source assembly 20 and thedetector assembly 24. After a prescribed time (e.g., 150 seconds)measured from the point of contrast injection has lapsed, the gantry 12then rotates about the breast 18 to generate two sets of image data(Step 204). In the illustrated embodiment, the two sets of image dataare generated in quick succession (e.g., within 5 to 20 milliseconds)using radiation at different levels. It should be noted that the timewithin which the first and the second sets of image data are generatedshould not be limited to the example, and that the first and the secondsets of image data can be generated within any time period as long asthe first and the second sets of image data are captured fast enough torender the object being imaged appear motionless. As the gantry 12rotates about the breast 18, the x-ray source assembly 20 alternatelyemits radiation at a first and a second energy levels. Particularly, theradiation should have a first energy level that is below a k-absorptionedge (K-edge) of the contrast agent, and a second energy level that isabove the k-edge of the contrast agent. The emitted radiation at bothlevels is attenuated by the breast 18 and impinges on the detectorassembly 24. The detector assembly 24, in turn, generates the first andthe second sets of image data in response to radiation at the first andsecond levels, respectively. Additional sets of image data for differentgantry angle can be generated as the gantry 12 rotates about the breast.After a desired number of sets of image data (e.g., sufficient forreconstruction of volumetric image) have been generated, the image datacan be stored in a computer readable medium for later processing. Insome embodiments, the gantry 12 makes at least one rotation to generatethe sets of image data. In alternative embodiments, the gantry 12 makesa partial rotation to generate the sets of image data.

In alternative embodiments, instead of acquiring the first and thesecond sets of image data during a gantry rotation, the first set ofimage data can be generated during a first rotation (or rotations), andthe second set of image data can be generated during a successiverotation (or rotations). In such case, the patient 16 can be instructedto hold breath while the gantry 12 is rotating about the breast 18 tocollect image data. This ensures or increases the chance that an objectcaptured at a position in the first set of image data will also becaptured at substantially the same position in the second set of imagedata. For example, to increase the chance that the image data taken atthe first and the second energy levels will have similar spatialregistration, the image data are taken at the end of an expiration withbreath holding. In other embodiments, a patient position monitoringsystem can be employed to monitor positions of the patient 16 as thegantry 12 is rotating about the breast 18. In such case, motion signalrepresentative of a physiological movement of the patient 16 can be usedto predictively gate an operation of the x-ray source assembly 20 suchthat image data can be generated at prescribed phase(s) or prescribedamplitude range(s) of a physiological cycle. In other embodiments, themotion signal can be synchronized with the image data to a common timebase, and CT volumetric images can be retrospectively constructed usingthe image data. Patient position monitoring systems, and systems andmethods for predictive gating have been described in U.S. patentapplication Ser. No. 09/893,122, filed Jun. 26, 2001, the entiredisclosure of which is hereby incorporated by reference.

After the image data for both radiation levels have been obtained, theimage data are then processed to create a composite image (Step 206). Insome embodiments, image data generated using radiation at the first andsecond energy levels are used to construct a first volumetric image anda second volumetric image, respectively. Various techniques can be usedto construct a volumetric image. Construction of volumetric images usingcone beam CT has been described in U.S. patent application Ser. No.10/656,438, entitled “Radiation process and apparatus”, filed Sep. 5,2003, which claims priority to U.S. Provisional Patent Application Ser.No. 60/416,022, the disclosures of which are expressly incorporated byreference herein. After the first and the second volumetric images areconstructed, they are processed to obtain the composite image. Forexample, the first volumetric image can be digitally subtracted from thesecond volumetric image to obtain a composite image. Alternatively, alogarithmic transform is applied to the first and the second volumetricimages, which converts values assigned to each pixel of the first andthe second volumetric image to a natural logarithm of the originalvalue. The first transformed image are then digitally subtracted fromthe second transformed image to obtain a composite image. Such techniqueremoves much of the dependence on the background breast thickness and anintensity of the x-ray exposure used from the subtracted image.Different scaling factors can also be applied to the first and thesecond volumetric images before the first volumetric image is subtractedfrom the second volumetric image. Other frame processing algorithms andtechniques can also be used to create the composite image. For example,a boxcar filter can be applied to the first and the second volumetricimages for noise reduction, as is known in the art. In alternativeembodiments, instead of creating the first and the second volumetricimages, the first image data and the second image data can be processedto obtain composite image data, and the composite image data are thenused to construct the composite image.

Various methods can be used to align the first image data with thesecond image data to create the composite image. In some embodiments,the first and the second image data are generated at the same gantryangles. For example, for the case in which the first and the second setsof image data are generated within a gantry rotation, the gantry 12 canstop rotating as the first and the second sets of image data are beinggenerated at a gantry angle. The gantry 12 then rotates to a nextposition to generate next sets of image data. In such case, since thefirst and the second sets of image data are generated at the same gantryangles, the first and the second volumetric images generated from thefirst and the second image data, respectively, would be automaticallyaligned with each other. For the case in which the first and the secondsets of image data are each generated at successive gantry rotations,the second image data can be generated (in a subsequent gantry rotation)at the same gantry angles at which the first image data are generatedduring a previous gantry rotation. In such case, since the first and thesecond sets of image data are also generated at the same gantry angles,the first and the second volumetric images generated from the first andthe second image data, respectively, would be automatically aligned witheach other.

In other embodiments, the first and the second image data are generatedat different gantry angles. In such cases, the first image data and thesecond image data are used to create a first volumetric image and asecond volumetric image, respectively. The first volumetric image canthen be aligned with the second volumetric image by translating and/orrotating either of the first and the second volumetric images relativeto the other, such that a feature in the first volumetric image matcheswith the same feature in the second volumetric image. Alternatively, anaverage shift in the gantry angle between the first image data and thesecond image data can be determined, and the average shift in the gantryangle can be used to align the first and the second volumetric images.In some embodiments, the average shift in the gantry angle between thefirst and second image data can also be taken into account whenconstructing either or both of the first and the second volumetricimages such that the first and the second volumetric images are aligned.In other embodiments, the first image data can be modified to align withthe second image data. For example, if first image data are generated atgantry angles=15° and 25° (measured from an arbitrary reference) usingradiation at the first energy level, and second image data are generatedat gantry angle=20° using radiation at the second energy level, then thefirst image data generated at gantry angles 15° and 25° can be processed(e.g., averaged, or interpolated) to obtain modified first image datathat correspond to a gantry angle of 20°. A preferred method foraccomplishing this interpolation is to use known techniques forinterpolation. For example, for band limited data, Fourier analysistechniques can be used to determine an exact interpolation. The modifiedfirst image data and the second image data can then be used to generatea first volumetric image and a second volumetric image, respectively. Insome embodiments, the shift in a gantry angle between the first imagedata (taken at first energy level) and the second image data (taken atsecond energy level) is selected to be small compared to a change ingantry angle between each image data that are taken at the same energylevel. Such technique will enhance image registration between the firstvolumetric image and the second volumetric image.

Those skilled in the art understand that a contrast agent having ak-edge attenuates radiation above and near its k-edge, and does notsubstantially attenuate radiation below and near its k-edge. Forexample, Iodine, which has a k-edge of approximately 33 keV, attenuatesradiation that is higher than about 33 keV, and does not substantiallyattenuate radiation below about 33 keV (FIG. 3A). As such, by performingframe subtraction of the two volumetric images (or variation of the twovolumetric images), non-contrast features, i.e., images of object thatdoes not contain contrast agent, are removed or reduced from thecomposite image, while contrast features, i.e., images of object thatcontains contrast agent, are enhanced. For example, subtraction of thefirst image from the second image will reduce an image contrast (makeless visible) for bone and tissue that do not contain the contrastagent, and retain or enhance an image contrast (make more visible) ofvessels that contain the contrast agent.

In some embodiments, the composite image can be used to identifycancerous tissue. Cancerous tissue or tumors may take up contrast agentfaster and to a greater degree than do normal tissue because of thedenser capillaries associated with tumor angiogenesis. As such, locationof potential cancerous tissue can be determined by identifying anyunusual concentration of the contrast agent represented in the compositeimage.

The above described imaging procedure has several advantages over thetraditional mammography. First, the use of a CT system eliminates theneed to perform breast compression, thereby eliminating any discomfortthat may be experienced by the patient 16 due to breast compression, andallowing transport of contrast agent into the breast 18. In addition,the above described imaging procedure eliminates superimposition ofstructures within the breast 18, and provides better contrast and detailresolution than those obtained from traditional mammography.

In the above described embodiment, a single type of contrast agent isused. However, in alternative embodiments, a plurality of contrastagents can be used. In such cases, each of the contrast agents can beselected such that tissue having a particular characteristic or featurecan be enhanced in a composite image. The plurality of contrast agentscan be simultaneously injected into the patient 16, or alternatively,administered to the patient 16 at different prescribed times.

Furthermore, instead of generating image data using radiation at a firstand a second energy levels, in alternative embodiments, image data canbe generated using radiation at more than two energy levels. In suchcases, the generated image data at different energy levels are processedto generate a composite image such that an appearance of a feature dueto a contrast agent can be enhanced or maximized in the contrast image.For example, image data using radiation at one or more energy levels canbe used to form a first set and a second set of integrated image data.Each of the first and second sets of integrated image data can be imagedata generated using radiation at one of the levels, or alternatively,can be created by integrating image data that have been generated usingradiation at a plurality of energy levels. The first set of integrateddata can then be subtracted from the second set of integrated image datato form a composite image, as similarly discussed previously.

In the above described embodiment, a volumetric composite image (orcomposite image data) is generated for a prescribed lapsed time afterthe injection of the contrast agent. However, in alternativeembodiments, steps 204 and 206 can be repeated to generate additionalvolumetric composite images (or composite image data) for differentlapsed time after the injection of the contrast agent. In such case,image data at the first and second radiation energy levels are generatedat different lapsed time after injection of the contrast agent.Volumetric images (or composite image data) associated with the firstand the second radiation energy levels can be constructed for each ofthe lapsed time using the image data, and the same technique describedpreviously can be used to obtain a composite image for each of thelapsed times. In some embodiments, the volumetric composite images (orcomposite image data) for the different lapsed times can be used todetermine time-resolved kinetics of the contrast agent. For example,features of an object within a region of interest in a composite imagecan be used to calculate an iodine concentration at a location of theobject (i.e., target site). After iodine concentrations that correspondto different times at which image data are generated have beendetermined, the iodine concentration for a target site and for a normaltissue site can be plotted versus time, and the graph for the targetsite can be compared with the graph for the normal tissue site todetermine whether the target site contains abnormal tissue. It is wellknown that kinetic curve for abnormal tissue can exhibit differentcharacteristic from the kinetic curve for a normal tissue.

FIG. 3B is an example of a graph 350 that can be generated using theabove described procedures. The graph 350 includes a first set of datapoints 352 representing concentration of the contrast agent at a targetsite that are generated at various times after the contrast agent hasbeen injected into the patient 16. The graph 350 also includes a secondset of data points 354 representing concentration of the contrast agentat a normal tissue region that are generated at various times afterinjection of the contrast agent. As shown in the graph 350, the datapoints 352, 354 show how concentration of the contrast agent increases(due to absorption of the contrast agent) and decreases (due toelimination of the contrast agent) over time, for the target site andthe normal tissue, respectively. Concentration of the contrast agent forabnormal tissue, such as tumor tissue or cancerous tissue, may have ahigher curve (or data points) due to additional vessel growths(angiogenesis) associated with the abnormal tissue.

X-Ray Source Assembly

As previously discussed, the x-ray source assembly 20 is configured togenerate radiation at a plurality of levels. The x-ray source assembly20 can be variously constructed to perform such function. FIG. 4A showsa x-ray source assembly 20 a in accordance with one embodiment of theinvention. The x-ray source assembly 20 a includes an electron gun 402(cathode), and a voltage supply 460 secured to the electron gun 402. Thevoltage supply 460 is configured to supply a voltage to the electron gun402 during use. The x-ray source assembly 20 a also includes a tube 410to which the electron gun 402 is secured, and a first target material474 and a second target material 476 (anodes) located inside a lumen 418of the tube 410. The target materials 474, 476 can include a variety ofmaterials that have suitable mechanical, thermal, electronic properties,and other suitable properties for production of prescribed x-ray spectraand intensity. Examples of materials that can be used includes holmium,erbium, lanthanum, cerium, praseodymium, neodymium, samarium, europium,terbium, dysprosium, thulium, ytterbium, lutetium, barium, molybdenum,rhodium, zirconium, hafnium, tungsten, titanium, rhenium, rhenium,molybdenum, copper, graphite, other rare earth materials and platinumgroup metals, and combination thereof. Suitably stable and refractorycompounds, such as cerium boride (CeB₆), and other compounds formed fromany of the above mentioned materials can be used for the targetmaterials 474, 476. In the illustrated embodiment, the target materials474, 476 are secured to a disk 416, and are positioned relative to eachother in a radial arrangement. The disk 416 is rotatably coupled to amotor 414. Alternatively, the disk 416 itself can be made from thetarget materials 474, 476. The target materials 474, 476 aresubstantially centered at an axis of rotation of the disk 416.Alternatively, either or both of the target materials 474, 476 can beoff-centered from the axis of rotation of the disk 416. The x-ray sourceassembly 20 a further includes an electron deflector 490 located withinthe lumen 418 of the tube 410. In the illustrated embodiment, theelectron deflector 490 comprises an electromagnetic field generator thatgenerates electromagnetic field, which changes a path of travelingelectrons as they exit the electron gun 402. The electron deflector 490is coupled to the processor 54 and/or the control 40, which controls anoperation of the electron deflector 490.

During use, the voltage supply 460 supplies a voltage to the electrongun 402 to generate a cloud of electrons. Due to a potential that isgenerated between the electron gun 402 and the first target material474, the electrons accelerate towards the first target material 474,forming a beam 440 of electrons. The beam 440 can be a continuous beam,or alternatively, a pulsed beam. X-rays at a first energy level aregenerated by the interaction of the electron beam 440 and the firsttarget material 474. Most of the generated x-rays are confined by thetube 410, while a beam 442 of the x-rays escape from an x-ray window420. As x-rays are generated, the motor 414 rotates the disk 416 suchthat the electron beam 440 impinges on different locations on the firsttarget material 474, thereby cooling the first target material 474. Togenerate x-ray beam at a second energy level, the processor 54 modulatesthe electron deflector 490 such that the electron deflector 490 deflectsexiting electrons from the electron gun 402 onto to the second targetmaterial 476. The deflected electrons impinge onto the second targetmaterial 476 and interact with the second target material 476 togenerate radiation at the second energy level, as similarly discussedpreviously. In one embodiment, the first generated radiation has anenergy level that is below a k-edge of a contrast agent, and the secondgenerated radiation has an energy level that is higher than a k-edge ofthe contrast agent.

In alternative embodiments, instead of having an electromagnetic fieldgenerator, the electron deflector 490 can have other configurations. Forexample, the electron deflector 490 can include a structure forphysically deflecting or selecting electrons emitted from the electrongun 402. In such case, the structure can be coupled to a positioner,which changes a position and/or orientation of the structure, therebydeflecting electrons towards different directions. Furthermore, inalternative embodiments, instead of having an electromagnetic fieldgenerator, the x-ray source assembly 20 a can include a magnetic fieldgenerator that generates magnetic field for deflecting electrons thatare accelerating towards either of the target materials 474, 476. Itshould be noted that the configuration (e.g., shapes, dimensions,designs, and arrangements of various components) of the x-ray sourceassembly 20 a should not be limited to the example illustrated in thefigure, and that the x-ray source assembly 20 a can have otherconfigurations. For example, in other embodiments, instead of securingthe target materials 474, 476 to a rotatable disk 416, the targetmaterials 474, 476 can be coupled to a reservoir of cooling fluid.

FIG. 4B shows another x-ray source assembly 20 b in accordance withother embodiments of the invention. The x-ray source assembly 20 b issimilar to the x-ray source assembly 20 a, except that it does not havethe electron deflector 490. In the illustrated embodiment, the first andthe second target materials 474, 476 are secured to the disk 416relative to each other in a circumferential arrangement (Section B-B).The motor 414 is coupled to the processor 54 and/or the control 40,which controls an operation of the motor 414. During use, the motor 414rotates the disk at a prescribed rate such that the first and the secondtarget materials 474, 476 can be alternately placed at a targetposition. Electrons emitted from the electron gun 402 impinges onto thetarget materials 474, 476 at the target position, and the interaction ofthe electrons with the target materials 474, 476 generate x-rayradiation at the first and the second levels, respectively. The speed ofrotation of the disk 416 can be modulated such that radiation at thefirst and second levels can be generated at prescribed times (e.g., insynchronization with image data collection). In some embodiments, theprocessor 54 controls an operation of the motor 414 such that generationof radiation at the first and second levels are synchronized withprescribed gantry angles. Although four first target materials 474 andfour second target materials 476 are shown, in alternative embodiments,the x-ray source assembly 20 b can have fewer or more than four sets ofthe first and the second target materials 474, 476. Also, in alternativeembodiments, a region of the disk 416 is not covered by a targetmaterial, thereby providing a time gap between successive radiationgeneration.

FIG. 4C shows another x-ray source assembly 20 c in accordance withother embodiments of the invention. The x-ray source assembly 20 c issimilar to the x-ray source assembly 20 a. However, instead of using anelectron deflector to modulate a direction of accelerating electrons,the x-ray source assembly 20 c includes a positioner 478 coupled to themotor 414. The positioner 478 is configured to move the disk 416 ineither of the directions as indicated by arrows 479. The positioner 478is coupled to the processor 54 and/or the control 40, which controls anoperation of the positioner 478. For example, in one embodiment, theprocessor 54 can gate a transmission of a signal to the positioner 478such that the positioner 478 can place either the first target material474 or the second target material 476 at a target position to which thebeam 440 of electron is directed. The placement of the first and thesecond target materials 474, 476 can be performed synchronously with agantry rotation. To generate radiation at a first energy level, thepositioner 478 places the first target material 474 at the targetposition. The voltage supply 460 then supplies a voltage to the electrongun 402 to generate the electron beam 440. X-rays at a first energylevel are generated by the interaction of the electron beam 440 and thefirst target material 474. To generate radiation at a second energylevel, the positioner 478 places the second target material 476 at thetarget position, and x-rays at the second energy level are generated bythe interaction of the electron beam 440 and the second target material476. In alternative embodiments, instead of positioning the targetmaterials 474, 476, a positioner (not shown) can be coupled to theelectron gun 402. In such case, the positioner is configured to positionthe electron gun 402 such that the electron beam 440 generated therefromwill either impinge the first target material 474 or the second targetmaterial 476.

FIG. 4D shows another x-ray source assembly 20 d in accordance withother embodiments of the invention. The x-ray source assembly 20 d issimilar to the x-ray source assembly 20 a except that it does not havean electron deflector and that it only has one target material 404. Thex-ray source assembly 20 d includes a first filter 464 and a secondfilter 466 secured to a positioner 462. The filters 464, 466 can be madefrom a variety of materials, such as aluminum, molybdenum, and copper.Also, any of the materials described previously for the target materialcan also be used. The preferred x-ray filter characteristics are suchthat the first filter 464 has a high x-ray transmission window in therange corresponding to a transmission window of the contrast materialjust below the contrast material k-edge, and the second filter 466 has atransmission window about the same width but above the contrast materialk-edge. For example, for Gd contrast agent, the filter materials for thefirst and the second filters 464, 466 may be selected from elements withatomic numbers below 64 (the atomic number of Gd) and above 64,respectively. When determining a preferred filter factors related to asource x-ray spectrum, filteration by tissue, and detector efficiencyspectral profile can be considered.

In the illustrated embodiment, the positioner 462 and the voltage supply460 are coupled to the processor 54 and/or the control 40, whichcontrols an operation of the positioner 462 and the voltage supply 460.For example, in one embodiment, the processor 54 can gate a transmissionof a signal to the positioner 462 such that the positioner 462 can placeeither the first filter 464 or the second filter 466 in front of thex-ray window 420 synchronously with a gantry rotation. To generateradiation having a first characteristic, the positioner 462 places thefirst filter 464 in front of the x-ray window 420. At least a portion ofthe x-rays generated from the interaction of the electron beam 440 andthe target material 404 exits the x-ray window 420 and impinges on thefirst filter 464. The impinging x-rays are filtered by the first filter464 to produce radiation having a first characteristic, e.g., a firstenergy level. To generate radiation having a second characteristic, thepositioner 462 replaces the first filter 464 in front of the x-raywindow 420 with the second filter 466. X-rays exiting the x-ray window420 impinges on the second filter 464, and are filtered by the secondfilter 464 to produce radiation having a second characteristic, e.g., asecond energy level. Those skilled in the art understand that radiationbeam generated using different filter material and/or filter thicknesswill have different characteristics. As such, different materials can beselected for construction of the filters 464, 466, and/or the thicknessof the filters 464, 466 can be designed such that x-ray radiationgenerated by the x-ray source assembly 20 c will have certain desiredcharacteristic(s).

It should be noted that in alternative embodiments, the x-ray sourceassembly 20 d can include only one filter. In such case, radiation atone of the first and the second energy levels can be generated using thefilter (i.e., applying a filter factor), and radiation at the other ofthe first and the second energy levels can be generated without usingany filter (i.e., applying a null filter factor). In addition, inalternative embodiments, the first and the second filters 464, 466 neednot be secured to a same structure. For example, the first filter 464can be secured to a first structure, and the second filter 466 can besecured to a second structure. In such cases, one or more motors can beused to alternately place the first and the second filters 464, 466 intothe generated x-ray radiation.

FIG. 4E shows another x-ray source assembly 20 e in accordance withother embodiments of the invention. The x-ray source assembly 20 e issimilar to the x-ray source assembly 20 d except that it does not havethe filters 464, 466. The x-ray source assembly 20 e includes a voltagesupply 452 having a switch 454 for switching a supplied voltage betweena first level and a second level. To generate radiation at a firstenergy level, the switch 454 causes the voltage supply 452 to generate avoltage having a first level such that x-ray 442 at the first energylevel can be generated. To generate radiation at a second energy level,the switch 454 causes the voltage supply 452 to generate a voltagehaving a second level such that x-ray 442 at the second energy level canbe generated. As similarly discussed previously, the voltage supply 452can be coupled to the processor 54 and/or the control 40. In someembodiments, the x-ray source assembly 20 e can further include one ormore filters (similar to the filters 464, 466) for generating radiationhaving certain desired characteristic(s).

FIG. 4F shows a x-ray source assembly 20 f in accordance with otherembodiments of the invention. The x-ray source assembly 20 f is similarto the x-ray source assembly 20 e except that it has two voltagesupplies 406, 408 secured to the electron gun 402. The first and thesecond voltage supplies 406, 408 are configured to supply a firstvoltage and a second voltage to the electron gun 402, respectively,during use. The voltage supplies 406, 408 are coupled to the processor54 and/or the control 40, which controls an operation of the voltagesupplies 406, 408. For example, the processor 54 can gate a transmissionof an activation signal to the voltage supplies 406, 408, such thatradiation at the first and second energy levels can be generated atprescribed gantry angles of rotation. During use, the first voltagesupply 406 supplies a first voltage to the electron gun 402 to generatea cloud of electrons. Due to a potential that is generated between theelectron gun 402 and the target material 404, the electrons acceleratetowards the target material 404, forming a beam 440 of electrons. X-raysat a first energy level are generated by the interaction of the electronbeam 440 and the target material 404. To generate x-ray beam at a secondenergy level, the first voltage supply 406 is deactivated and the secondvoltage supply 408 is activated to supply a second voltage to theelectron gun 402. In one embodiment, the first voltage supply 406supplies a first voltage for generating radiation having an energy levelthat is below a k-edge of a contrast agent, and the second voltagesupply 408 supplies a second voltage for generating radiation having anenergy level that is higher than a k-edge of the contrast agent. Forexample, the first and the second voltages can be selected such that thefirst and the second x-ray range is between 25% to 50% below and above,respectively, a k-edge of the contrast agent.

It should be noted that although several examples of the x-ray sourceassembly 20 have been described, the scope of the invention should notbe so limited. In alternative embodiments, the x-ray source assembly 20can have other configurations as long as the x-ray source assembly 20can deliver radiation at a plurality of energy levels. In addition, inalternative embodiments, a feature described in reference to anembodiment of the x-ray source assembly 20 can be combined with otherembodiment(s) of the x-ray source assembly 20. For example, in analternative embodiment, the x-ray source assembly 20 can include twovoltage supplies and two filters. In another embodiment, the x-raysource assembly 20 can have two target materials and two voltagesupplies or two filters. Furthermore, in alternative embodiments, thex-ray source assembly 20 can be configured to generate radiation at morethan two energy levels. For examples, the x-ray source assembly 20 caninclude more than two target materials, more than two filters, and/ormore than two voltage supplies. Other x-ray source assembly capable ofgenerating radiation at different energy level can also be used.Radiation sources capable of generating X-ray radiation at differentenergy levels are described in U.S. patent application Ser. No.10/033,327, entitled “RADIOTHERAPY APPARATUS EQUIPPED WITH ANARTICULABLE GANTRY FOR POSITIONING AN IMAGING UNIT”, filed on Nov. 2,2001, the entirety of which is expressly incorporated herein byreference.

Detector Assembly

As discussed previously, the detector assembly 24 generates imagesignal/data in response to radiation impinges thereon. The detectorassembly 24 can be variously constructed. FIG. 5 shows a detectorassembly 24 a in accordance with some embodiments of the invention. Thedetector assembly 24 a comprises an imager 500 that includes a x-rayconversion layer 502 made from a scintillator element, such as CesiumIodide (CsI), and a photo detector array 504 (e.g., a photodiode layer)coupled to the x-ray conversion layer 502. The x-ray conversion layer502 generates light photons in response to x-ray radiation, and thephoto detector array 504, which includes a plurality of detectorelements 506, is configured to generate electrical signal in response tothe light photons from the x-ray conversion layer 502. In theillustrated embodiment, both the x-ray conversion layer 502 and thephoto detector array 504 are pixilated, thereby forming a plurality ofimaging elements 508. However, the x-ray conversion layer 502 may benon-pixilated in an alternative embodiment. In the illustratedembodiment, the imager 500 has a curvilinear surface (e.g., a partialcircular arc). Such configuration is beneficial in that each of theimaging elements 508 of the imager 500 is located substantially the samedistance from the x-ray source 20 assembly. In alternative embodiments,the imager 500 can have a rectilinear surface or a surface having otherprofiles. In the illustrated embodiment, each image element 508 (orpixel) has a cross sectional dimension that is approximately 200 micronsor more, and more preferably, approximately 400 microns or more.However, image elements having other dimensions may also be used.Preferred pixel size can be determined by a prescribed spatialresolution. The image elements 508 having 200 to 400 microns in crosssectional dimension are good for general anatomy imaging. For breastimaging, image elements 508 having cross sectional dimension that isbetween 50 to 100 microns are preferred. The imager 500 can be made fromamorphous silicon, crystal and silicon wafers, crystal and siliconsubstrate, or flexible substrate (e.g., plastic), and may be constructedusing flat panel technologies (e.g., active-matrix flat paneltechnologies) or other techniques known in the art of making imagingdevice.

In one embodiment, each of the image elements 508 comprises a photodiode(forming part of the detector element 506) that generates an electricalsignal in response to a light input. The photodiode receives light inputfrom the x-ray conversion layer 502 that generates light in response tox-rays. The photodiodes are connected to an array bias voltage to supplya reverse bias voltage for the image elements. A transistor (such as athin-film N-type FET) functions as a switching element for the imageelement 508. When it is desired to capture image data from the imageelements 508, control signals are sent to a gate driver to “select” thegate(s) of transistors. Electrical signals from the photodiodes“selected” by the gate driver are then sent to charge amplifiers, whichoutputs image signals/data for further image processing/display.

In one embodiment, the image data are sampled from the image elements508 one line at a time. Alternatively, the image data from a pluralityof lines of the image elements 508 can be sampled simultaneously. Sucharrangement reduces the time it takes to readout signals from all linesof image elements 508 in the imager 500. This in turn, improves a framerate (i.e., number of frames that can be generated by the imager 500 persecond) of the imager 500. Devices and methods for simultaneouslycollecting image data from a plurality of lines of image elements havebeen described in U.S. patent application Ser. No. ______, entitled“Multi-slice flat panel computed tomography”, filed concurrentlyherewith, the entire disclosure of which is expressly incorporated byreference herein.

During use, radiation at a first energy level impinges on the detectorassembly 24 a, which then generates image signals/data in response tothe radiation at the first energy level. After the image signals/dataare read out from the photo detector array 504, radiation at a secondenergy level is directed to the detector assembly 24 a. The assembly 24a then generates image signals/data in response to the radiation at thesecond energy level. In one embodiment, one or more filters can beplaced between the x-ray source assembly 20 and the detector assembly 24(e.g., on top of the conversion layer 502) before radiation at either orboth of the energy levels is directed to the detector assembly 24 a. Thefilter(s) alters radiation exiting from the patient 16, such thatradiation having a desired characteristic will be received by thedetector assembly 24 a. In one embodiment, a first filter(s) can be usedto maximize or optimize a detective quantum efficiency of the detectorassembly 24 a for radiation at a first energy level, while a secondfilter(s) can be used to maximize or optimize detective quantumefficiency of the detector assembly 24 a for radiation at a secondenergy level. For example, the detector assembly 24 a may have a uniformsensitivity to all photon energies in a spectrum, may have a sensitivitythat is proportional to photon energy, or may have “holes” where photonsof certain energy ranges are not efficiently absorbed. For each of thesedifferent types of detector assembly 24 a, one or more filters can beselected to maximize an efficiency of the system 10 (e.g., maximizing aresponse of the system 10 in measuring the injected contrast agent,and/or minimizing dose delivery and time). The placement of thefilter(s) can be accomplished manually or mechanically. In someembodiments, the filters can be parts of the detector assembly 24.

In alternative embodiments, the detector assembly 24 may use differentdetection schemes. For example, in alternative embodiments, instead ofhaving the x-ray conversion layer 502, the detector assembly 24 caninclude an imager having a photoconductor, which generateselectron-hole-pairs or charges in response to x-ray. FIG. 6schematically shows an imager 600 constructed in accordance withalternative embodiments of the present invention. The imager 600, whichcan be a flat panel imager, for example, includes a x-ray conversionpanel 610 aligned with a detector array 620. The x-ray conversion panel610 includes a first electrode 602, a second electrode 604, and aphotoconductor 606 secured between the first electrode 602 and thesecond electrode 604. The electrodes 602, 604 can be made from a widevariety of materials, such as silver, chromium, aluminum, gold, nickel,vanadium, zinc, palladium, platinum, carbon, etc., and alloys of thesematerials. The photoconductor 506 can be made from a variety of photoconductive materials, such as mercuric Iodide (HgI₂), Lead Iodide(PbI₂), Bismuth Iodide (BiI₃), Cadmium Zinc Telluride (CdZnTe),Amorphous Selenium (a-Se), or equivalent thereof. HgI₂ and PbI₂ areparticularly preferred because these materials efficiently absorb x-rayphotons and have desirable photo conductive properties. Photoconductorscreens (or panels) from these materials can increase a modulationtransformer function (MTF) value—a measure of spatial resolution,thereby providing high radiograph quality. Other materials known in theart may also be used. The photoconductor 606 may be a single orpoly-crystalline layer, or an amorphous layer. The photoconductor 606 ispreferably deposited by physical vapor depositon (PVD) or particle inbinder process (PIB). Alternatively, the photoconductor 606 may also besecured to the first and second electrodes 602 and 604 by a suitableadhesive, depending on the materials from which the photoconductor 606and the first and second electrodes 602 and 604 are made. Othertechniques known in the art may also be used to secure thephotoconductor 606 to the first and second electrodes 602 and 604.Photoconductors and imagers made therefrom are well known in the art,and therefore would not be described in further details herein.

When using the imager 600, the first and second electrodes 602 and 604are biased by a voltage source to create a potential difference or abias between the first and second electrodes 602 and 604. The biasedelectrodes 602 and 604 create an electric field across the regionbetween the first and second electrodes 602 and 604. When thephotoconductor 606 is irradiated by x-ray, a response, such as electronhole pairs (EHPs) or charges, are generated and drift apart under theinfluence of the electric field across the region between the first andsecond electrodes 602 and 604. The charges are collected by the detectorarray 620, which includes a plurality of detector elements 622 arrangedin a two-dimensional array. The detector elements 622 are configured togenerate electric signals in response to the charges collected on thefirst electrode 602. In one embodiment, the detector elements 622 areamorphous silicon (a-Si:H) charge detectors. Each detector element 622may have a storage capacitor to store the charge generated by the X-raysand collected by the first electrode 602. Each detector element 622 mayalso include a switching element, such as a thin film transistor (TFT),a switching diode, or the like, to access the collected charge byreadout electoronics. Optionally the detector elements 622 can containfurther components for signal or charge buffering and amplification. Thedetector elements 622 can also include polycrystalline silicon ororganic active elements. Each of the detector elements 622 forms a pixelof the X-ray image generated using the detector array 620. The detectorarray 620 also includes a pixel access circuit (not shown) coupled todetector elements 622. The pixel access circuit accesses the detectorelements 622 and reads the electric signals from the detectors elements622. In one embodiment, pixel access circuit includes a gate driver thatgenerates row access signals to sequentially access detector elements622 by rows and reads electric signals out of detector elements 622 bycolumns. Each row access signal can access either a single row ormultiple rows of detectors elements 622. Likewise, each read action canread electric signals from either a single column or a plurality ofcolumns of the detectors elements 622. The process of accessing detectorelements 622 and reading electric signals there from is well known inthe art, and therefore, would not be describe in further detail. In someembodiments, one or more filters can be placed between the x-ray sourceassembly 20 and the detector assembly 24 (e.g., on top of the electrode602 or on top of the photoconductor 606) before radiation at either orboth of the energy levels is directed to the detector assembly 24, assimilarly discussed previously.

Other detection schemes can also be used. In alternative embodiments,the detector assembly 24 can be configured to detect photon pulseamplitude and/or photon count. Such arrangement allows a pulse amplitudespectrum of one or more x-ray photon events to be measured on a pixel bypixel basis. In such case, the pulse amplitudes can be processed tocreate desired image data. For example, in one embodiment, image datacan be created by considering pulse amplitudes that are above aprescribed threshold. Detectors capable of detecting photon pulseamplitudes and photon count have been described in U.S. patentapplication Ser. No. 10/438,684, the entire disclosure of which isexpressly incorporated by reference herein.

FIG. 7 shows another imager 700 in accordance with other embodiments ofthe invention. The imager 700 includes a plurality of first imagingelements 702 and second imaging elements 704. Each of the first and thesecond imaging elements 702, 704 is similar to the imaging element 508described previously.

In the illustrated embodiments, each of the first imaging elements 702includes a first conversion element (or a scintillating material) havinga first radiation conversion characteristic, and each of the secondimaging elements 704 includes a second conversion element having asecond radiation conversion characteristic. By means of non-limitingexamples, a radiation conversion characteristic can be a sensitivity ofreaction to radiation, a quantity of photons created per unit ofradiation, a photon generation efficiency, and other variables relatedto any of these characteristics. For example, the first conversionelement can be made from one of the materials selected from the groupthat includes mercuric Iodide (HgI₂), Lead Iodide (PbI₂), Bismuth Iodide(BiI₃), Cadmium Zinc Telluride (CdZnTe), and Amorphous Selenium (a-Se),while the second conversion element can be made from another materialselected from the group. Each of these materials has a different k-edge,and the screen thickness can be chosen to generate a detector with“holes” (low efficiency bands) and “sinks” (high efficiency bands) thatare below and above the k-edge(s), respectively. In some embodiments,either or both of the first and the second conversion elements can bemade from more than one materials. This allows multiple “holes” and“sinks” at various k-edge related photon energies be generated. Thefirst and the second conversion elements together form a conversionpanel. The imager 700 also includes a photo detector array aligned withthe conversion panel. The photo detector array comprises a plurality ofdetector elements configured to generate a signal in response to lightphotons received from the conversion panel. An access circuit (notshown) is coupled to the photo detector array and is configured tocollect signals from one or more lines of the detector elements in thephoto detector array. In some embodiments, all of the detector elementshas similar functional characteristics. In other embodiments, the photodetector array can include a plurality of first detector elementsconfigured to generate signals in response to photons having a firstenergy level, and a plurality of second detector elements configured togenerate signals in response to photons having a second energy level.

In alternative embodiments, each of the first and the second imagingelements 702, 704 includes a photoconductor layer, as similarlydiscussed previously. In such case, each of the first imaging elements702 includes a first photoconductor element (e.g., a layer) having afirst charge generation characteristic, and each of the second imagingelements 704 includes a second photoconductor element having a secondcharge generation characteristic. By means of non-limiting examples, acharge generation characteristic can be a sensitivity of reaction toradiation, a quantum level of charges created per unit of radiation, acharge generation efficiency, and other variables related to any ofthese characteristics. Any of the materials discussed previously withreference to FIG. 6 can be used to construct the first and the secondphotoconductor elements. For example, the first photoconductor elementcan be made from one material, and the second photoconductor element canbe made from another material. Alternatively, or additionally, the firstand the second photoconductor elements can also have differentthicknesses. The first and the second photoconductor elements togetherform a photoconductor layer. The imager 700 also includes a photodetector array aligned with the photoconductor layer. The photo detectorarray comprises a plurality of detector elements configured to generatea signal in response to charges received from the photoconductor layer.An access circuit (not shown) is coupled to the photo detector array andis configured to collect signals from one or more lines of the detectorelements in the photo detector array. In some embodiments, all of thedetector elements has similar functional characteristics. In otherembodiments, the photo detector array can include a plurality of firstdetector elements configured to generate signals in response to chargeshaving a first quantum level, and a plurality of second detectorelements configured to generate signals in response to charges having asecond quantum level.

In the illustrated embodiment, the first and the second image elements702, 704 are arranged in rows (or columns), such that a row of the firstimage elements 702 are located adjacent a row of the second imageelements 704. In alternative embodiments, instead of having thealternate row arrangement, the first and the second image elements 702,704 can be positioned relative to each other in other arrangements. Forexample, the first and the second image elements 702, 704 can bearranged relative to each other in a checkerboard pattern.

In some embodiments, the first and the second image elements 702, 704are used to generate image signal/data in response to radiation at afirst energy level and a second energy level, respectively. Missing databetween rows of the first image elements 702, and missing data betweenrows of the second image elements 704 can be generated by interpolation.In other embodiments, the first image elements 702 only generate imagesignal/data in response to radiation at a second energy level, and aredeactivated when radiation at a first energy level is being applied. Thesecond image elements 704 are activated when radiation at the first andsecond energy levels are being applied. In such case, missing databetween rows of the first image elements 702 can be interpolated toobtain a first set of image data for a first image that corresponds toradiation at the second energy level. To obtain a second image thatcorresponds to radiation at the first energy level, missing data betweenrows of the second image elements 704 can be interpolated to obtain asecond set of image data, and the first set of image data is thensubtracted from the second set of image data to generate the secondimage. Other similar techniques or algorithms can also be used.

It should be noted that the configuration of the x-ray imager 24 shouldnot be limited to the examples discussed previously. By way of example,U.S. patent application Ser. No. 10/439,350, entitled “MULTI ENERGYX-RAY IMAGER” filed on May 15, 2003, discloses x-ray imaging devicescapable of generating signals in response to multiple radiation energylevels, and can be used as the detector assembly 24. U.S. patentapplication Ser. No. 10/439,350 is incorporated herein by reference inits entirety. In addition, U.S. patent application Ser. No. 10/013,199entitled “X-RAY IMAGE ACQUISITION APPARATUS” and filed on Nov. 2, 2001discloses an X-ray image detecting device that is capable of detectingmultiple energy level X-ray images, and can also be used as the detectorassembly 24 in accordance with the present invention. U.S. patentapplication Ser. No. 10/013,199 is incorporated herein by reference inits entirety.

In the above described embodiments, the detector assembly 24 includes asingle imager for generating image data. However, in alternativeembodiments, the detector assembly 24 can include a plurality ofimagers, with each configured to generate image data in response to aprescribed range of radiation levels. FIG. 8 shows a detector assembly24 b that includes a plurality of imagers, in accordance with someembodiments of the invention. The detector assembly 24 b includes afirst detector 800, and a second detector 802 located behind the firstdetector 800. The first detector 800 is configured to generate imagesignal/data in response to radiation at a first level, and the seconddetector 800 is configured to generate image signal/data in response toradiation at a second level. Either or both of the detectors 800, 802can include a layer of scintillating material or a photoconductor, assimilarly discussed previously. During use, radiation at a first energylevel impinges on the first detector 800 and the first detector 800generates a first set of image signal/data in response to the radiation.After the image signal/data are read from the first detector 800,radiation at a second energy level is directed to the detector assembly24 b. In some embodiments, radiation at the second energy level is notsubstantially attenuated by the first detector 800, thereby allowing thesecond detector 802 to generate image signal/data in response toradiation that passes through the first detector 800. Alternatively, thefirst detector 800 can be constructed such that radiation can passbetween image elements of the first detector 800. For example, each ofthe image elements of the first detector 800 can be separated from anadjacent image element such that a gap is provided to allow radiation topass to the second detector 802. In such case, the first detector 800generates image signal/data in response to radiation at a first energylevel impinging thereon, and the second detector 802 generates imagesignal/data in response to radiation at a second energy level that haspassed between the image elements of the first detector 800.

Computed Tomography (CT) and Cone Beam CT Reconstruction

CT or CBCT defines a volume, Vol_(point), around an image point andmeasures the average x-ray attenuation of the material in this volumerelative to that of water. In rectilinear (parallel beam) geometry thisvolume is independent of location of the point in the image. Theachievement of good reconstructed image quality for other geometriesdepends on the success of the reconstruction algorithm's ability toachieve this independence for other data collection geometries.

Denoting S_(Obj)(E_(eff), r) as the average object attenuation at apoint in an object being imaged, and averaging over the reconstructionpoint volume Vol_(point) where s_(Obj)(E, r) is the objects attenuationcross-section, and r_(Obj)(r) is the objects density at position r,yields:

${\Sigma_{Obj}( {E_{{eff}\; \_ \; v},r} )} \equiv \frac{\int_{0}^{{Vol}_{point}}{{{\sigma_{Obj}( {E,r} )} \cdot \rho_{Obj}}{r}}}{\int_{0}^{{Vol}_{point}}{1{r}}}$

where it is assumed that the photon energy spectrum is considered adelta function and that the E_(eff) _(—) _(v) is chosen appropriately.For a uniform material such as water this becomes: Σ_(H2O)(E_(eff) _(—)_(v), r)≡σ_(H2O)(E)·ρ_(H2O). Hence, if a point in the object has anattenuation HU_(Object), the Vol_(point) averaged attenuation at thatpoint,

${S_{Obj}( E_{{eff}\; \_ \; v} )},{{{is}\text{:}{\Sigma_{Obj}( E_{{eff}\; \_ \; v} )}} \equiv {\frac{( {{HU}_{Object} + 1000} )}{1000} \cdot ( {{\sigma_{H\; 20}( E_{{eff}\; \_ \; v} )} \cdot \rho_{H\; 20}} )}}$

An A-to-D counts (ADCounts) as measured by a detector, for the case whenno object is placed between a radiation source and the detector assemblycan be expressed in the following equation:

ADCountsThru_(Air) ⋅ Δ A ≡ (S_Fact ⋅ ∫φ(E) ⋅ ψ det (E)E) ⋅ Δ A

where φ (E) represents a flux (x-ray photons incident per unit area)created by a radiation source, as determined by source characteristics,and ψdet(E) represents a detector efficiency. ΔA should be less than orequal to the nominal area cross section of the reconstruction volume. AnADCounts for the case when an object is placed between the radiationsource and the detector assembly can be expressed by the followingequation:

${{{{ADCountsThru}_{Obj}(L)} \cdot \Delta}\; A} \equiv {\begin{bmatrix}{{S\_ Fact} \cdot {\int{{{\varphi (E)} \cdot \psi}\; {{\det (E)} \cdot}}}} \\{^{- {({\sum\limits_{L = 0}^{ThruObject}{{{\Sigma_{Obj}{({E_{{eff}\; \_ \; v},r})}} \cdot \Delta}\; L}})}}{E}}\end{bmatrix} \cdot \text{?}}$?indicates text missing or illegible when filed

where S_Fact is a scaling factor that represents voltage per photon, andE_(eff) _(—) _(v) represents effective energy that has taken intoconsideration a change of sigma (Σ_(Obj)).

In the above equation, a change in ψ (E) along a path is neglected. Achange in φ (E) with position has an affect on S_(Obj)(E_(eff), r). Thereconstruction algorithm itself assumes that the S_(Obj)(E_(eff), r) isindependent of the direction of the beam defining the L direction andalso the direction of the incident beam passing along. Generally, as afirst approximation, the above considerations are neglected, and it isassumed that the photon energy spectrum is considered a delta functionand that the E_(eff) _(—) _(s) is chosen appropriately to consider achange of spectrum of the flux along a path. Appropriate E_(eff) _(—)_(v) and E_(eff) _(—) _(s) are not necessarily equal. Then the abovegives the transmission as follows:

${{TransmissionAlong}( {L,E_{{eff}\; \_ \; s}} )} \equiv {\frac{{{{ADCountsThru}_{Obj}(L)} \cdot \Delta}\; A}{{{ADCountsThru}_{Air} \cdot \Delta}\; A}\mspace{14mu} {{or}\lbrack {{{\sum\limits_{L = 0}^{ThruObject}( {{{\Sigma_{Obj}( {E_{{eff}\; \_ \; {sv}},r} )} \cdot \Delta}\; L} \rbrack} \equiv {- {\ln( \frac{{ADCountsThru}_{Obj}(L)}{{ADCountsThru}_{Air}} )}}},} }}$

where E_(eff) _(—) _(sv) represents effective energy that has taken intoconsideration the change of sigma and a change of a spectrum of the fluxalong a path. CBCT uses these line integrals to reconstruct the value ofΣ_(Obj)(E_(eff) _(—) _(v), r) throughout the object volume. The effectof the various approximations discussed above can be reduced byreconstructing the difference image between a uniform “norm” objectconstructed from material e.g. H₂O which results in similar observedline integrals. Then,

${{{{{ADCountsThru}_{H\; 20}(L)} \cdot \Delta}\; A} \equiv \begin{bmatrix}{{S\_ Fact} \cdot {\int{{{\Phi (E)} \cdot \psi}\mspace{11mu} {{\det (E)} \cdot}}}} \\{^{- {({\sum\limits_{L = 0}^{{ThruH}\; 20}{{{\Sigma_{H\; 20}{({E_{{eff}\; \_ \; v},r})}} \cdot \Delta}\; L}})}}{E}}\end{bmatrix}}{{\cdot \Delta}\; A}$ ${{and}\begin{bmatrix}{\sum\limits_{L = 0}^{ThruObject}( {{\Sigma_{Obj}( {E_{{eff}\; \_ \; {sv}},r} )} -} } \\{{ {\Sigma_{H\; 20}( {E_{{eff}\; \_ \; {sv}},r} )} ) \cdot \Delta}\; L}\end{bmatrix}} \equiv {- {{\ln ( \frac{{ADCountsThru}_{{Obj}^{(L)}}}{{ADCountsThru}_{H\; 20}} )}.}}$

The quantity Σ_(Obj)(E_(eff) _(—) _(sv), r)−Σ_(H2O)(E_(eff) _(—) _(sv),r) can then be reconstructed.

Similarly, when using a filter in the incident beam, a change in thebeam energy spectrum can be represented as follows:

$\lbrack {\sum\limits_{L = 0}^{ThruObject}{{\begin{pmatrix}{{\Sigma_{Obj}( {E_{{eff}\; \_ \; {sov}},r} )} -} \\{\Sigma_{Obj}( {E_{{eff}\; \_ \; {sfv}},r} )}\end{pmatrix} \cdot \Delta}\; L}} \rbrack \equiv {- {\ln ( \frac{{ADCountsSpectNoFilter}_{{Obj}^{(L)}}}{{ADCountsSpectFiltered}_{{Obj}^{(L)}}} )}}$

where E_(eff) _(—) _(sov) represents effective energy when no filter isused, and E_(eff) _(—) _(sfv) represents effective energy when a filteris used. In such case, the quantity Σ_(Obj)(E_(eff) _(—) _(sov),r)−Σ_(H2O)(E_(eff) _(—) _(sfv), r) can be reconstructed. Those skilledin the art understand that similar techniques can be used for the caseswhen different contrast agents, different voltages and/or differenttarget materials are used to create radiation at multiple energy levels.Other reconstruction algorithms known in the art can also be used. Conebeam CT has been described in U.S. patent application, entitled, “Amulti-mode cone beam CT radiotherapy simulator and treatment machinewith a flat panel imager”, filed Dec. 17, 2002, the entire disclosure isexpressly incorporated by reference herein.

Computer System Architecture

FIG. 9 is a block diagram that illustrates an embodiment of a computersystem 900 upon which an embodiment of the invention may be implemented.Computer system 900 includes a bus 902 or other communication mechanismfor communicating information, and a processor 904 coupled with the bus902 for processing information. The processor 904 may be an example ofthe processor 54, or alternatively, an example of a component of theprocessor 54, of FIG. 1. The computer system 900 also includes a mainmemory 906, such as a random access memory (RAM) or other dynamicstorage device, coupled to the bus 902 for storing information andinstructions to be executed by the processor 904. The main memory 906also may be used for storing temporary variables or other intermediateinformation during execution of instructions to be executed by theprocessor 904. The computer system 900 further includes a read onlymemory (ROM) 908 or other static storage device coupled to the bus 902for storing static information and instructions for the processor 904. Adata storage device 910, such as a magnetic disk or optical disk, isprovided and coupled to the bus 902 for storing information andinstructions.

The computer system 900 may be coupled via the bus 902 to a display 912,such as a cathode ray tube (CRT), for displaying information to a user.An input device 914, including alphanumeric and other keys, is coupledto the bus 902 for communicating information and command selections toprocessor 904. Another type of user input device is cursor control 916,such as a mouse, a trackball, or cursor direction keys for communicatingdirection information and command selections to processor 904 and forcontrolling cursor movement on display 912. This input device typicallyhas two degrees of freedom in two axes, a first axis (e.g., x) and asecond axis (e.g., y), that allows the device to specify positions in aplane.

The invention is related to the use of computer system 900 forcollecting and processing image data. According to one embodiment of theinvention, such use is provided by computer system 900 in response toprocessor 904 executing one or more sequences of one or moreinstructions contained in the main memory 906. Such instructions may beread into the main memory 906 from another computer-readable medium,such as storage device 910. Execution of the sequences of instructionscontained in the main memory 906 causes the processor 904 to perform theprocess steps described herein. One or more processors in amulti-processing arrangement may also be employed to execute thesequences of instructions contained in the main memory 906. Inalternative embodiments, hard-wired circuitry may be used in place of orin combination with software instructions to implement the invention.Thus, embodiments of the invention are not limited to any specificcombination of hardware circuitry and software.

The term “computer-readable medium” as used herein refers to any mediumthat participates in providing instructions to the processor 904 forexecution. Such a medium may take many forms, including but not limitedto, non-volatile media, volatile media, and transmission media.Non-volatile media includes, for example, optical or magnetic disks,such as the storage device 910. Volatile media includes dynamic memory,such as the main memory 906. Transmission media includes coaxial cables,copper wire and fiber optics, including the wires that comprise the bus902. Transmission media can also take the form of acoustic or lightwaves, such as those generated during radio wave and infrared datacommunications.

Common forms of computer-readable media include, for example, a floppydisk, a flexible disk, hard disk, magnetic tape, or any other magneticmedium, a CD-ROM, any other optical medium, punch cards, paper tape, anyother physical medium with patterns of holes, a RAM, a PROM, and EPROM,a FLASH-EPROM, any other memory chip or cartridge, a carrier wave asdescribed hereinafter, or any other medium from which a computer canread.

Various forms of computer-readable media may be involved in carrying oneor more sequences of one or more instructions to the processor 904 forexecution. For example, the instructions may initially be carried on amagnetic disk of a remote computer. The remote computer can load theinstructions into its dynamic memory and send the instructions over atelephone line using a modem. A modem local to the computer system 900can receive the data on the telephone line and use an infraredtransmitter to convert the data to an infrared signal. An infrareddetector coupled to the bus 902 can receive the data carried in theinfrared signal and place the data on the bus 902. The bus 902 carriesthe data to the main memory 906, from which the processor 904 retrievesand executes the instructions. The instructions received by the mainmemory 906 may optionally be stored on the storage device 910 eitherbefore or after execution by the processor 904.

The computer system 900 also includes a communication interface 918coupled to the bus 902. The communication interface 918 provides atwo-way data communication coupling to a network link 920 that isconnected to a local network 922. For example, the communicationinterface 918 may be an integrated services digital network (ISDN) cardor a modem to provide a data communication connection to a correspondingtype of telephone line. As another example, the communication interface918 may be a local area network (LAN) card to provide a datacommunication connection to a compatible LAN. Wireless links may also beimplemented. In any such implementation, the communication interface 918sends and receives electrical, electromagnetic or optical signals thatcarry data streams representing various types of information.

The network link 920 typically provides data communication through oneor more networks to other devices. For example, the network link 920 mayprovide a connection through local network 922 to a host computer 924 orto a medical equipment 926. The data streams transported over thenetwork link 920 can comprise electrical, electromagnetic or opticalsignals. The signals through the various networks and the signals on thenetwork link 920 and through the communication interface 918, whichcarry data to and from the computer system 900, are exemplary forms ofcarrier waves transporting the information. The computer system 900 cansend messages and receive data, including program code, through thenetwork(s), the network link 920, and the communication interface 918.

Although the above embodiments have been described with reference todual-energy, contrast-enhanced cone beam CT imaging of breasts, thescope of the invention should not be so limited. In alternativeembodiments, any of the above described devices or methods (or similardevices or methods) can be used to perform imaging of other portions ofa body, such as a liver, a chest, a heart, or other vascular structures.In addition, besides using the above techniques for detecting cancerousbreast tissue, any of the devices and/or methods described herein canalso be used to detect other cancerous tissue or other types of tissue.For example, a contrast agent and/or energy levels of radiation can beselected such that feature(s) of tissue having certain characteristicscan be enhanced using any of the above described techniques.

Also, in alternative embodiments, instead of generating a compositeimage for static visualization, a plurality of composite images can becreated for different phases (or phase intervals) of a physiologicalcycle, and the created composite images can be displayed in a sequenceto form a video. For example, a patient positioning monitoring systemcan be used to collect motion data representative of a motion of thepatient 16, while the gantry 12 rotates about the patient 16 (or objectbeing imaged) to generate image data using radiation at a first andsecond energy levels. The motion data and the image data can beretrospectively synchronized to a common time base, thereby allowingcomposite images that correspond to different times of the physiologicalcycle to be created. In one embodiment, the image data can betime-binned based on prescribed phase ranges of a physiological cycle.In an alternative embodiment, the image data can be time-binned based onprescribed amplitude ranges of the physiological cycle. Systems andmethods for monitoring patient's position, time-binning based on phaseor amplitude ranges, and retrospective gating have been describe in U.S.patent application Ser. No. ______, entitled, “Method and system forradiation application”, filed on Oct. 3, 2003, the entire disclosure ofwhich is herein incorporated by reference.

In addition, instead of using cone beam CT, other imaging techniques,such as Spiral CT, Fan Beam CT, laminar tomography, MRI, or PET, can beused in a similar process to obtain a composite image from a first imageand a second image. For example, for MRI, a contrast agent that affectsan environment of protons can be selected.

Furthermore, besides using the system 10 to obtain dual-energycontrast-enhanced images, any of the embodiments of the system 10, orcomponents of the system 10, can be used for other applications. Forexample, in some embodiments, the x-ray source assembly 20 can deliverradiation at a first level for imaging a target object, and radiation ata second level for treating the target object.

Although particular embodiments of the present inventions have beenshown and described, it will be understood that it is not intended tolimit the present inventions to the preferred embodiments, and it willbe obvious to those skilled in the art that various changes andmodifications may be made without departing from the spirit and scope ofthe present inventions. For example, the operations performed by theprocessor 54 can be performed by any combination of hardware andsoftware within the scope of the invention, and should not be limited toparticular embodiments comprising a particular definition of“processor”. In addition, the term “image” as used in this specificationincludes image data that may be stored in a circuitry or acomputer-readable medium, and should not be limited to image data thatis displayed visually. The specification and drawings are, accordingly,to be regarded in an illustrative rather than restrictive sense. Thepresent inventions are intended to cover alternatives, modifications,and equivalents, which may be included within the spirit and scope ofthe present inventions as defined by the claims.

1-34. (canceled)
 35. A method for generating image data, comprising:applying a first voltage to generate radiation at a first energy level;generating a first set of image data based at least in part on theradiation at the first energy level; applying a second voltage togenerate radiation at a second energy level; generating a second set ofimage data based at least in part on the radiation at the second energylevel; and creating composite image data using the first and the secondsets of image data; wherein the first and the second voltages areapplied using a single voltage supply.
 36. The method of claim 35,wherein the composite image data is created by subtracting the first setof image data from the second set of image data.
 37. The method of claim35, wherein the creating the composite image data comprises: modifyingthe first set of image data; modifying the second set of image data; andsubtracting the first modified set of image data from the secondmodified set of image data.
 38. The method of claim 37, wherein thesteps of modifying comprises applying a logarithmic transform to thefirst and the second sets of image data. 39-42. (canceled)
 43. Themethod of claim 35, wherein the first voltage is applied to create afirst potential using a target material for generating the radiation atthe first energy level, and the second voltage is applied to create asecond potential using the target material for generating the radiationat the second energy level.
 44. The method of claim 43, wherein thefirst potential is created between an electron gun and the targetmaterial, and the second potential is created between the electron gunand the target material, the first potential being different from thesecond potential.
 45. The method of claim 35, wherein the first voltageand the second voltage are applied to an electron gun.
 46. The method ofclaim 35, wherein the radiation at the first energy level is generatedusing a first filter, and the radiation at the second energy level isgenerated using a second filter that is different from the first filter.47. The method of claim 46, wherein the first filter comprises a nullfilter.
 48. The method of claim 35, wherein the first voltage is appliedto create a first potential using a first target material for generatingthe radiation at the first energy level, and the second voltage isapplied to create a second potential using a second target material forgenerating the radiation at the second energy level, the first targetmaterial being different from the second target material.
 49. A systemfor generating image data, comprising: a voltage supply configured forapplying a first voltage to generate radiation at a first energy level,and for applying a second voltage to generate radiation at a secondenergy level; an imager for generating a first set of image data basedat least in part on the radiation at the first energy level, and forgenerating a second set of image data based at least in part on theradiation at the second energy level; and a processor for creatingcomposite image data using the first and the second sets of image data.50. The system of claim 49, wherein the processor is configured tocreate the composite image data by subtracting the first set of imagedata from the second set of image data.
 51. The system of claim 49,wherein the processor is configured for creating the composite imagedata by: modifying the first set of image data; modifying the second setof image data; and subtracting the first modified set of image data fromthe second modified set of image data.
 52. The system of claim 49,wherein the processor is configured to perform the acts of modifying byapplying a logarithmic transform to the first and the second sets ofimage data.
 53. The system of claim 49, further comprising: a targetmaterial; wherein the voltage supply is configured to apply the firstvoltage to create a first potential using the target material forgenerating the radiation at the first energy level, and the voltagesupply is configured to apply the second voltage to create a secondpotential using the target material for generating the radiation at thesecond energy level.
 54. The system of claim 53, further comprising: anelectron gun; wherein the first potential is created between theelectron gun and the target material, and the second potential iscreated between the electron gun and the target material, the firstpotential being different from the second potential.
 55. The system ofclaim 49, further comprising an electron gun, wherein the voltage supplyis configured to apply the first voltage and the second voltage to theelectron gun.
 56. The system of claim 49, further comprising: a firstfilter; and a second filter that is different from the first filter;wherein the radiation at the first energy level is generated using thefirst filter, and the radiation at the second energy level is generatedusing the second filter.
 57. The system of claim 56, wherein the firstfilter comprises a null filter.
 58. The system of claim 49, furthercomprising: a first target material; and a second target material thatis different from the first target material; wherein the voltage supplyis configured to apply the first voltage to create a first potentialusing a first target material for generating the radiation at the firstenergy level, and the voltage supply is configured to apply the secondvoltage to create a second potential using a second target material forgenerating the radiation at the second energy level.